The incidence of infections after total joint replacement surgery has increased over the past decade despite the widespread use of intravenous antibiotic prophylaxis and a focus on aseptic surgical technique. Post-arthroplasty infections still occur in about 1.2% of primary arthroplasties and 3-5% of revisions. As the demand for joint replacements increases with the aging population, the total number of infections is projected to rise from 17,000 to 266,000 per year by 2030 as the number of arthroplasties exceeds 3.8 million surgeries. The treatment of a post-arthroplasty infection is exceedingly difficult. Bacteria (especially S. aureus) form extracellular anionic polysaccharide biofilms on implanted metallic/plastic materials that block penetration of immune cells and antibiotics, promoting bacterial survival. Once a biofilm is formed, surgical removal of all the implanted materials is necessary. Most of these infections are caused by staphylococcal species (about 70%) and an increasing number are due to virulent antibiotic-resistant strains such as methicillin-resistant S. aureus (MRSA), which further complicate treatment.
The current standard of care in the U.S. to treat a chronic post-arthroplasty infection is a two-stage procedure beginning with (1) surgical removal of all prosthetic components and bone cement, debridement of necrotic/granulation tissue, placement of an antibiotic-impregnated spacer, administration of a 6-week course of intravenous antibiotics (during which the patient is unable to bear weight on the affected limb), and (2) revision arthroplasty after the infection has cleared. In severe infections and refractory cases, arthrodesis, resection arthroplasty and amputation are sometimes necessary. In the elderly, these infections result in increased mortality. Overall, the treatment of post-arthroplasty infection involves extensive medical and surgical care, prolonged disability/rehabilitation and significantly worse outcomes. In addition, these infections represent an enormous economic burden due to additional medical costs and resource utilization as well as indirectly through lost wages and productivity. These medical costs alone average $144,514 (compared with $30,173 for an uncomplicated arthroplasty), which correspond to an annual national healthcare burden of $8.63 billion by 2015.
Most post-arthroplasty infections are thought to be caused by invading bacteria at the time of surgery. As treatment of infected implanted materials is exceedingly difficult, especially due to the inherent difficulties in treating an established biofilm, one potential therapeutic strategy is to focus on the prevention of infection.
One way to avoid infection is to use implantable devices that deliver a drug, such as an antibiotic, directly to the implantation site. Local delivery of certain drugs can be more effective than traditional systemic routs, as certain tissues, particularly bone tissue, have limited vascularity. Additionally, local delivery allows for a high local concentration while avoiding systemic side-effects.
Local delivery of a large bolus dose at the time of surgery would not provide long term effects. While pumps to deliver drugs to a local site may be used in certain cases, they are not feasible in all circumstances and can be cumbersome.
In order to achieve local, continuous delivery of a drug, medical devices can be coated with a drug in a manner that would allow the sustained and localized release of the drug.
Implantable medical devices can be made from various materials, including, but not limited to, metals, polymers or a combination of different materials. Metals commonly used in implantable medical devices include, but are not limited to, titanium and stainless steel. Common polymers, include, but are not limited to, polyethylene and polypropylene. However, due to the differences in surface energies between polymers and metals, what may be a suitable coating on one material will not be effective on another.
While metals have surface energies of around 100, the surface energy of a polymer is typically around 30. The relative surface energies of a surface and coating material affect the ability of the coating material to effectively adhere to the surface. In order for a liquid (such as a coating solution) to optimally adhere to a surface, it must thoroughly “wet out” the surface to which it is to be bonded. “Wetting out” means that the liquid flows and covers a surface to maximize the contact area and the attractive forces between the liquid and solid surface. For a liquid such as an adhesive or coating solution to effectively wet out a surface, the surface energy of the liquid must be as low as or lower than the surface energy of the substrate. Standard adhesive or coating formulations wet out and bond to high surface energy (HSE) surfaces such as metal or ABS plastic, but fail to bond to low surface energy (LSE) polyolefins that include polypropylene and polyethylene.
For traditional structural adhesives or coatings to bond low surface energy substrates such as polyolefins, surface treatments, such as exposure to UV light or treatment with chromic acid, have been used to raise the substrate surface energy by as much as 30% to better meet the adhesive surface energy. Other strategies to modify the surface properties and precisely tune interfacial interactions of materials include, lithographic patterning, binary assembly, anodic oxidation, electrodeposition and chemical etching, plasma etching, laser treating, ion bombardment, UV light inducement, surfactants, chemical oxidation treatment, polymer modification, electrospinning, electrochemical etching, chemical vapor deposition, sol-gel processing, and so on. Although high quality surfaces can be fabricated by the above mentioned approaches, these methods all have some disadvantages limiting their further applications, such as the complexity of experimental setup, rigorous preparation conditions, higher energy cost and the dependence on the specific surface chemistries. Moreover, these methods are only suitable for some given substrates and cannot be applied to a wide range of surfaces or substrates.
For example, current state of the art drug-eluting stents usually have one to three or more layers in the coating e.g. a base layer for adhesion, a main layer for holding the drug, and sometimes a top coat to slow down the release of the drug and extend its effect. For example, the CYPHER® stent requires an initial base-layer of parylene to allow for adhesion of the drug containing polymer. Replacing these multiple coats with a single coating would result in more straightforward manufacturing.
US Patent Application No. 2007/0198040 A1 describes a bioresorbable polymer coating on a surgical mesh as a carrier for the antimicrobial agents rifampin and minocyline. However, a coating suitable for a polypropylene mesh may not provide enough adhesion to medical devices, such as orthopedic pins, which are made of metal and/or undergo significant manipulation and abrasion during surgical installation.
Therefore, there is a need for polymeric coatings that can provide improved adhesion to substrates with varying surface properties. Furthermore, these coatings should also be biocompatible in order to avoid rejection; sturdy/sticky to avoid peeling off during implantation; biodegradable/resorbable so there is no long term foreign body response; capable of sustained delivery of drugs; easily tailored to deliver variety of drugs; easily tailored for coating onto different substrates; easily applied to a variety of devices by spraying; dipping or melting; and compatible with other excipients. It has been surprisingly discovered that a blend of certain polyphenolic polymers and polyethylene glycol does provide such properties when coated onto medical devices having a wide range of different surface characteristics.